Magnetic resonance imaging (MRI) is a medical imaging modality that can create images of the inside of a human body without using x-rays or other ionizing radiation. MRI uses a powerful magnet to create a strong, uniform, static magnetic field (i.e., the “main magnetic field”). When a human body, or part of a human body, is placed in the main magnetic field, the nuclear spins that are associated with the hydrogen nuclei in tissue water or fat become polarized. This means that the magnetic moments that are associated with these spins become preferentially aligned along the direction of the main magnetic field, resulting in a small net tissue magnetization along that axis (the “z axis,” by convention). An MRI system also comprises components called gradient coils that produce smaller amplitude, spatially varying magnetic fields when a current is applied to them. Typically, gradient coils are designed to produce a magnetic field component that is aligned along the z axis and that varies linearly in amplitude with position along one of the x, y or z axes. The effect of a gradient coil is to create a small ramp on the magnetic field strength and concomitantly on the resonant frequency of the nuclear spins, along a single axis. Three gradient coils with orthogonal axes are used to “spatially encode” the MR signal by creating a signature resonance frequency at each location in the body. Radio frequency (RF) coils are used to create pulses of RF energy at or near the resonance frequency of the hydrogen nuclei. The RF coils are used to add energy to the nuclear spin system in a controlled fashion. As the nuclear spins then relax back to their rest energy state, they give up energy in the form of an RF signal. This signal is detected by the MRI system and is transformed into an image using a computer and known reconstruction algorithms.
During a patient scan, the gradient coil(s) of the gradient coil assembly that produce the magnetic field gradients dissipate large amounts of heat. A cooling system or apparatus may be provided to remove the heat generated by the gradient coils. The maximum attainable performance of a gradient coil may be limited by the heat removal capability of the cooling system or systems used. With modern high power MRI imaging sequences, it is increasingly difficult to remove the larger levels of heat generated by the gradient coils. In addition, imposed limits on the temperature rise inside a gradient coil and on the resulting temperature elevation in the patient bore during scanning may result in duty cycle limitations for aggressive imaging sequences.
The gradient coil assembly used in an MRI system may be a shielded gradient coil assembly that consists of inner and outer gradient coil assemblies bonded together with a material such as epoxy resin. Typically, the inner gradient coil assembly includes inner (or main) coils of X-, Y- and Z-gradient coil pairs or sets and the outer gradient coil assembly includes the respective outer (or shielding) coils of the X-, Y- and Z-gradient coil pairs or sets. The Z-gradient coils are typically cylindrical with a conductor spirally wound around the cylindrical surface (or mandrel). The transverse X- and Y-gradient coils are commonly formed from a copper panel with an insulating backing layer. A conductor turn pattern (e.g., a fingerprint pattern) may be cut in the copper layer which leaves an interturn spacing (or gap) between the adjacent turns.
One prior cooling system includes using a hollow conductor for the main Z-gradient coil located on the outside of the inner gradient coil assembly. A cooling fluid may then be passed through the hollow Z-conductor to remove heat generated by the gradient coil assembly. Such a cooling arrangement, however, does not provide direct cooling to the transverse X- and Y-gradient coils of the inner gradient coil assembly. Typically, heat generated by the transverse X- and Y-gradient coils must be transferred to the hollow Z-gradient coil via insulation layers between the gradient coils of the inner gradient coil assembly. The heat transfer may be limited, therefore, by the thermal resistance of the insulation layer material (e.g., glass-epoxy resin).
In order to provide additional cooling, a dedicated cooling channel or channels may be provided on, for example, an inside diameter of the inner gradient coil assembly. Another option for providing additional cooling is using a hollow conductor for the transverse X-gradient coil of the inner gradient coil assembly. Additional dedicated cooling channels or additional hollow inner gradient coils, however, increase the radial space of the gradient coil assembly and also move the main X-, Y- and Z-gradient coils to positions at larger radii. These may result in reduced gradient strength for the gradient coil assembly.
There is a need, therefore, for a gradient coil cooling system that allows for the removal of heat and for the control of the temperature rise inside the gradient coil without the addition of radial space to the gradient coil assembly.